1. Field of the Invention
The present invention relates to an optical tomographic imaging apparatus, and more particularly, to an optical tomographic imaging apparatus that is used for ophthalmological care, tomographic observation of skin, tomography scan of a digestive and cardiovascular wall with an endoscope or a catheter constituted of the optical tomographic imaging apparatus, or the like.
2. Description of the Related Art
In recent years, an optical interference tomographic imaging method and an optical interference tomographic imaging apparatus, to which a low coherence interferometer technology or a white light interferometer technology is applied, are in actual use.
In particular, an optical tomographic imaging apparatus (optical interference tomographic imaging apparatus) that performs an optical coherence tomography (OCT) utilizing interference phenomenon of multi-wavelength light may be used to obtain a tomographic image of a sample with high resolution.
Therefore, in the ophthalmological field, an optical tomographic imaging apparatus is becoming an indispensable apparatus for obtaining a tomographic image of a fundus or a retina.
In addition to the ophthalmological application, the optical tomographic imaging apparatus has also been used for tomographic observation of skin, tomography scan of a digestive and cardiovascular wall with an endoscope or a catheter constituted of the apparatus, or the like. Hereinafter, the optical tomographic imaging apparatus is referred to as an OCT apparatus.
When imaging a living organism, a disturbance of the image due to a motion of the living organism (so-called a motion artifact) becomes a problem with the OCT apparatus in various application. In particular, in the OCT apparatus for the ophthalmological image diagnosis, existence of an eye movement largely affects accuracy of diagnosis.
As a typical eye movement, a motion of approximately 100 μm per second occurs both in the in-plane direction of the fundus (hereinafter referred to as a horizontal direction) and in the depth direction thereof (hereinafter referred to as a vertical direction) in a three-dimensional manner.
Therefore, by an OCT apparatus of time-domain method, which was first put into practical use in the ophthalmological application, a three-dimensional image technically could not be obtained because it requires a long period of time for imaging.
In this method, the imaging time is up to approximately one second for one B scan cross section image (two-dimensional image including one-dimensional image of the horizontal direction and one-dimensional image of the vertical direction). Therefore, a relatively long period of time is necessary for obtaining approximately 100 shots of B scan cross section images that are necessary for obtaining the three-dimensional image, and hence it is not practical because the motion artifact occurring due to the eye movement during the period is large.
For this reason, high speed performance of the OCT apparatus has further been desired.
In recent years, an OCT apparatus of a Fourier domain method (hereinafter referred to as an FD-OCT apparatus) has been widespread in use for its high speed performance that is at least ten times as fast as that of the conventional time-domain method.
Next, a schematic structure of this FD-OCT apparatus is described.
FIG. 10A is a schematic diagram of a typical FD-OCT apparatus for an ophthalmological use.
In FIG. 10A, a light beams emitted from a light source 1001 is guided by a single mode optical fiber 1002 and enters a fiber optical coupler 1003.
The fiber optical coupler 1003 is a so-called 2×2 type, which splits the incident light from the fiber 1002 to be caused to enter two output fibers.
One of the output fibers is coupled to an imaging optical system for human fundus that is a signaling beams path of a Michelson interferometer, and the other output fiber is coupled to a reference beams path of the interferometer.
In the signaling beams path, the light output from a fiber end is converted into a parallel beams by a collimate lens 1004, propagates in space, and enters an XY scanner 1005.
The XY scanner 1005 is a reflection type optical scanning apparatus that performs two-dimensional reflection angle control, and hence a reflected signaling beams is guided by a scanning lens 1006 and an ocular lens 1007 so as to enter a human eye 1008.
The XY scanner, the scanning lens and the ocular lens constitute a scanning optical system, which focuses the signaling beams as the parallel beams onto a fundus observation target region 1009 together with an optical action of the eye, and the focus position scans a surface of the fundus that is substantially perpendicular to the optical axis in a two-dimensional manner.
The ocular lens 1007 works to adjust the focus position in the depth direction. Control for scanning and focus is performed by a controlling and signal processing device 1101 that is connected to the XY scanner 1005 and a focus driving actuator 1010, in an integrative manner including other control.
A reflection beams from the fundus observation target region 1009 and a signaling beams propagating backward in substantially the same optical path among the backscattered light beams pass through the collimate lens 1004 again and returns to the fiber optical coupler 1003.
On the other hand, the reference beams is split by the fiber optical coupler 1003, is converted into a parallel beams by the collimate lens 1004, and is reflected by a reference beams mirror 1011 disposed on an optical delay driving apparatus 1012 so as to propagate backward along the optical path.
The position of the reference beams mirror 1011 is adjusted and controlled by controlling the optical delay driving apparatus 1012 together with, in particular, correction of an axial length that is different among individuals so that a total optical path length of the reference beams path becomes a predetermined length with the signaling beams path as a reference.
A translational stage including the reference beams mirror 1011 is connected to the controlling and signal processing device 1101 and is controlled together with other control in an integrative manner.
The reference beams propagating backward passes through the collimate lens 1004 again and returns to the fiber optical coupler 1003.
The signaling beams and the reference beams which have returned to the fiber optical coupler 1003 are split individually into components returning to the light source 1001 and components directed to an interfering beams receiving system. The signaling beams and the reference beams propagate in the same single mode fiber, i.e., are superimposed with each other so as to cause optical interference.
The interfering beams receiving system is a spectroscope in this example of the conventional technique, and the OCT apparatus constitutes a so-called spectral domain OCT apparatus (hereinafter referred to as an SD-OCT).
The interfering beams is converted into a parallel beams by the collimate lens 1004 and guided to a diffraction grating 1014 by a reflecting mirror 1013, and an action of the diffraction grating causes a first order diffraction light of the interfering beams to be directed to different angles according to a wavelength component contained in the same.
The individual wavelength components of the interfering beams that enter an imaging lens 1015 at different angles are focused for imaging at different positions on a line sensor 1016 according to the angles, and are read out as light intensities corresponding to individual pixels of the line sensor so that a signal thereof is sent to the controlling and signal processing device 1101.
Next, a structure and an action of the controlling and signal processing device 1101 are described with reference to FIG. 10B.
The controlling and signal processing device 1101 controls the XY scanner 1005, the optical delay driving apparatus 1012, the focus driving actuator 1010 and the line sensor 1016, and includes drivers and an acquisition unit for acquiring signals sent after detecting the angle, the position and the optical signal. Among the signals, a line image acquisition unit 1107 receives a light intensity signal train transmitted from the line sensor, and an FFT processing unit 1108 performs inverse fast Fourier transform on the signal train, and hence a result of the process is sent to a central processing unit 1103.
The central processing unit 1103 receives a digital optical interference signal sent after the inverse Fourier transform in time series and compares the digital optical interference signal with the following signals.
The digital optical interference signal is compared with a scanner position signal and a synchronizing signal from an XY scanner driver 1102, a delay position signal and a synchronizing signal from an optical delay driver 1105, and a focus position signal from a focus driver 1106.
Thus, the optical interference signal is associated with a position on the fundus observation target region.
After that, the optical interference signal is assigned to each of predetermined pixels, and hence the image is formed and displayed on an image displaying unit 1104.
Such an FD-OCT apparatus enables three-dimensional measurement of a fundus in an imaging time of approximately 1 to 3 seconds.
On the other hand, with regard to the OCT apparatus for ophthalmological use, an OCT apparatus having higher performance is demanded for early detection of diabetic retinopathy, glaucoma and age-related macular degeneration that are three major diseases that can cause loss of sight.
Specifically, an OCT apparatus having high resolution is demanded for detecting a minute change of a lesion in early stage.
An object to be imaged and measured is, for example, a change in an optic nerve fiber, a photoreceptor cell or a microvessel.
Among the resolutions, a vertical resolution, i.e., a resolution in the depth direction depends on characteristics of the light source used for the OCT apparatus. Therefore, the OCT apparatus has been devised to enlarge a wavelength width of light from the light source.
On the other hand, a horizontal resolution has a trade-off relationship with an optical spot size and a depth of focus. Therefore, simply constituting a focusing optical system having high numerical aperture (NA) is not sufficient.
The above-mentioned point is described below concretely with reference to equations and numerical examples.
The resolution of the OCT apparatus can be discussed as two resolutions in the cross section direction (vertical direction) and in the horizontal direction that is perpendicular to the cross section.
Among the two resolutions, the resolution in the cross section direction is determined by a wavelength width of light from the light source. As the wavelength width is larger, the resolution in the cross section direction is higher.
In other words, a narrow range in the vertical direction is rendered. The vertical resolution (Rz) is inversely proportional to a wavelength width of light from the light source, or in a strict sense, a wavelength width Δλ that is detected by the system after receiving light from the light source. The vertical resolution (Rz) is expressed by (Equation 1) below.Rz=kz×(λ^2/Δλ)  (Equation 1)where kz represents a constant that is approximately 0.4.
In a practical OCT for ophthalmological use, Δλ has been improved up to approximately 30 to 50 nm, and currently up to approximately 100 nm, while the corresponding vertical resolution is approximately 3 μm, which is becoming close to a modification in cell level described above.
On the other hand, the resolution (Rxy) in the horizontal direction is determined by an optical imaging resolution.
In other words, the resolution (Rxy) in the horizontal direction is determined by a numerical aperture (NA) of the imaging system and accompanying optical aberration.
Supposing that there is no aberration, the horizontal resolution is expressed by (Equation 2) below.Rxy=k1×(λ/NA)  (Equation 2)where k1 is a constant that is approximately 0.5.
On the other hand, a depth of focus (DOF) of the imaging system is expressed by (Equation 3) below.DOF=k2×(λ/NA^2)where k2 is a constant that is approximately 0.6.
In other words, high horizontal resolution and large depth of focus have a trade-off relationship based on an optical principle. For instance, if the horizontal resolution is doubled and the diameter of the optical spot size is halved, the depth of focus becomes one fourth as being inversely proportional to the square.
In a fundus diagnosis apparatus in which the OCT is most practical, numerical values of λ=0.84 μm and NA=0.02 approximately are used, for example. If these numerical examples and the above-mentioned (Equation 2) and (Equation 3) are used, there are derived Rxy=20 μm and DOF=2 mm approximately.
The thickness of a retina of a human eye is approximately 0.5 to 1 mm. For easiness of measurement and for avoiding a deviation from the imaging range due to various movements, an imaging range of approximately 2 mm in the depth direction is usually secured.
This is derived as a DOF value, and therefore the horizontal resolution is controlled to be 20 μm at most as the diameter of the optical spot size. This value of resolution is low by approximately one digit compared with 3 μm of the vertical resolution, but it is difficult to obtain a higher horizontal resolution with a simple structure.
In contrast, Japanese Patent Application Laid-Open No. 2007-101250 discloses a zone focusing OCT apparatus, in which multiple focus zones of high NA optical system having a small DOF are set, and images split in the depth direction are recombined, to thereby obtain high horizontal resolution over a wide range of depth of focus.
Such a zone focusing can be achieved by driving the focusing lens to be at multiple focus positions, focusing in a sequential manner while performing the imaging process, and recombining the images split in the depth direction.
In addition, Japanese Patent Application Laid-Open No. 2007-54251 discloses a method of calculating and setting a drive position of the focusing lens based on a specific position as a reference.